Magnetic resonance imaging (MRI) is now in widespread commercial practice. Although magnetic resonance spectroscopic imaging (MRSI) has long been envisioned as a useful further improvement, it is still largely in the experimental laboratory and/or prototype stage of development. In part, difficulties in achieving commercially practical MRSI include the fact that only relatively small NMR RF signal responses are available from other than hydrogen nuclei in a typical animal body environment (which is typically the sort of object to be imaged). While MRI is itself a sophisticated procedure requiring careful control of many physical parameters (e.g., spatial distribution of magnetic and radio fields), MRSI requires a substantially greater degree of care and precision.
For example, in typical MRI as presently practiced, a variation in the nominally static magnetic field B.sub.o of as much as about 30 parts per million (ppm) can easily be tolerated. In fact, although commercially available cryogenic magnets for use in magnetic resonance imaging typically include a set of 12 magnetic field "shim" coils (to compensate for inhomogeneities), there is often no need to even use the shim coils for conventional MRI --over even relatively large image volumes. (It should be observed, however, that the automated shimming techniques of this invention may also find application to improve even MRI.)
However, for MRSI, the homogeneity of the nominally static magnetic field must be held to considerably greater tolerances for acceptable results. For example, where the human head is involved, we desire a total variation of static magnetic field B.sub.o over a one liter image volume of no more than 0.2 ppm (and preferably on the order of no more than about 0.1 ppm).
In MRSI, attainable signal-to-noise ratio is determined by global homogeneity over the entire image volume--even though spectral resolution is determined only by local homogeneity over a given voxel. In the past, even when global homogeneity of 1.0 ppm has been attempted using available manual techniques, it typically required an experienced NMR scientist to expend on the order of 0.5 to 1.5 hours to achieve even this level of homogeneity. Since differences in magnetic susceptibility from one patient to the next can easily cause unacceptable changes in static magnetic field homogeneity within the image volume, some sort of shimming procedure therefore becomes necessary as part of the initial setup overhead time before conventional MRSI imaging procedures can be effected.
Since subsequent MRSI imaging procedures can easily require 30 minutes or so, it can be seen that available iterative manual techniques for achieving shim current adjustments would make the overall imaging procedure time so long as to be intolerable both from the standpoint of patient endurance and any economic measure of patient throughput per machine-day. Furthermore, before MRSI can achieve widespread commercial acceptance, the procedures must be sufficiently simple and reliable to be confidently accomplished by lower level technicians rather than only by high level NMR scientists.
Typically, in commercially available static magnet structures suitable for this field, there are plural (typically twelve) so-called "shim" coils in addition to the main magnetic field producing structure (whether it be a cryogenic, resistive or permanent magnet). Each of these "shim" coils has its own non-uniform spatial distribution of magnetic field component (typically of interest only along the z-axis direction). Accordingly, by adjusting the current flow in a given shim coil, a non-homogeneous component can be added to the existing field. If the added non-uniform shimming components are properly selected, then they may largely cancel unavoidable non-uniformities included in the nominally static field produced by the main field producing structure. Since there are typically many degrees of freedom (e.g., twelve) and since the coil fields are typically not all orthogonal with respect to one another (i.e., adjusting one coil to improve one area of non-uniformity may require readjustment of other shim coil currents so as to maintain earlier achieved compensations), it can be appreciated that the typical manual iterative process that is involved in "shimming" can become a very frustrating and time consuming project in its own right.
This problem has been recognized in the art and some solutions have been proposed. In our opinion, of presently known proposals by others, the following are thought to be especially relevant to our invention.
Prammer et al, "A New Approach to Automatic Shimming", J. Mag. Res., Vol. 77, pp 40-52 (1988). PA1 U.S. Pat. No. 4,680,551--O'Donnell et al. (1987).
Based on comments in Prammer et al (which is considered to be more relevant than O'Donnell et al.), it appears that further references 3 and 5-7 noted in that article may also be relevant to this invention.
O'Donnell et al directly measure actual magnetic field strength with a positionable probe at the outer surface of an imaginary sphere which encloses an image volume. These actual measurements outside the image volume, are used to infer magnetic field strengths within the image volume and they then use a weighted mean-square calculation algorithm for determining shim coil currents required to minimize inhomogeneity of the main magnetic field. It should be noted that the actual measurements of magnetic field are made in the absence of all shimming fields. Their technique does not appear suitable for use with an actual patient in place--and, indeed, they only report corrected inhomogeneities on the order of plus/minus 50 ppm which are clearly unsuitable for MRSI. Rather, it appears that O'Donnell et al were probably directing their efforts towards one-time factory calibration of shim current coils for traditional MRI.
Prammer et al., on the other hand, do specifically direct their attention to automatic shimming for improvement of chemical shift imaging. They use a modified Fourier imaging technique to quickly derive a simple phase measurement for each voxel to be shimmed. In effect, the phase difference between two spin echoes having slightly different time delays is measured--during which time delay nuclear phase shifts proportional to the static field are accumulated. Accordingly, the difference between the measured phases should be a function of the actual magnetic field--and thus should permit mapping of field inhomogeneities. Taking this as a given, Prammer et al. then use previously acquired differential field distribution data for each of the shim coils in conjunction with an appropriate calculation algorithm (Chebychev or Least-Square criterion) to calculate adjusted shim coil currents for reducing or minimizing resultant composite static magnetic field variations within the image volume. It should be noted that the field mapping method of Prammer et al requires a readout gradient much stronger than the largest field gradient of the unshimmed field.
Although the Prammer et al field mapping technique by use of a relatively simple and quickly acquired phase data set theoretically supplies the required information in a fairly short time (they report about 1 minute per plane), even Prammer et al. specifically note that their results produce apparent field differences that are inherently indistinguishable from chemical shift frequency changes produced by different NMR species. For example, the NMR response from human tissue typically includes water and fat as two major components of NMR species. Although Prammer et al. specifically state that "precautions must be taken for separating data points where lipid signal dominates over the water peak," they offer no suggestion as to how this is to be accomplished. Furthermore, they report only data taken with respect to a phantom having uniform density and distribution of NMR species.
Although Prammer et al. specifically recognize that it has long been known that chemical shift imaging may be used to map the inhomogeneity of the static magnetic field (citing to Maudsley), Prammer et al. specifically teach away from such mapping for automatic shimming purposes as it is "inherently slow, because it is inefficient to collect a spectrum at each spatial point in order to derive a single scalar value (the magnitude of the local field)."